Traditionally, X-ray diagnostic processes record x-ray image patterns on silver halide films. These systems typically direct an initially uniform pattern of impinging X-ray radiation, or X-rays, through an object to be studied, intercept the modulated pattern of X-ray radiation after the X-rays pass through the object to be studied with an X-ray radiation intensifying screen, record the modified pattern on a silver halide film, and chemically transform the latent pattern into a permanent and visible image called a radiograph.
Radiographs are produced by using layers of radiation sensitive materials to directly capture radiographic images as modulated patterns of electrical charges. Depending on the intensity of the incident X-ray radiation, electrical charges generated, either electrically or optically (indirectly via a scintillator) from the X-ray radiation, are quantized using a regularly arranged array of discrete solid-state radiation sensors.
Recently, there has been rapid development in the area of large area, flat-panel, digital X-ray imagers for digital radiology using active matrix technologies. An active matrix includes a two-dimensional array (of which, each element is called a pixel) of thin-film-transistors (TFTs) made with a large area compatible semiconductor material.
There are two general approaches to manufacturing flat-panel X-ray imagers. These approaches can be seen as a direct approach or an indirect approach. The direct approach primarily uses an amorphous selenium photoconductor as the X-ray to electric charge converting layer coupled directly to the active matrix. In the indirect approach, a phosphor screen or scintillator (e.g. CsI, GdOS etc) is used to convert X-rays to light photons which are then converted to electric charge using an additional pixel level light sensor fabricated with the TFT on the active matrix array.
Although the concept of large area dual energy X-ray imaging by stacking two X-ray imagers on top of each other has been around even in the era of computed radiography, the difficulty in acquiring a good quality image, increase in the overall time of the procedure and also the diagnostic time associated with two images, has limited its use in mainstream medicine. Stacking two imagers also drives the cost of production up linearly. More recently, with the rise of digital X-ray imagers, a technique called fast kVp switching has gained some acceptance where the X-ray source is switched within typically <100 ms between the high and low X-ray exposures while using a single digital imager to acquire the images in sequence. While a dual energy image can be obtained using this approach, this technique is susceptible to motion artifacts for example, due to patient breathing. Also, the digital imager can only be optimized to capture either the low or high-energy spectrum but this leads to a tradeoff between resolution and dose efficiency.
Commercially available flat-panel imagers are optimized to sense a single spectrum of incident X-ray energies e.g. a chest radiography imager is optimized to sense a 100 kVp spectrum or alternately, a mammography imager is designed to sense a 30 kVp spectrum. The optimization is undertaken usually by selecting an optimum thickness of the photoconductor for direct imaging or alternately, an optimum thickness of a scintillator for indirect imaging. If the scintillator is too thin, absorption efficiency suffers and if the scintillator is too thick, blurring increases.
Accordingly, there is a need for an improved digital X-ray imager.